Lu1-xI3:Cex - a scintillator for gamma ray spectroscopy and time-of-flight PET

ABSTRACT

The present invention concerns very fast scintillator materials comprising lutetium iodide doped with Cerium (Lu 1-x I 3 :Ce x ; LuI 3 :Ce). The LuI 3  scintillator material has surprisingly good characteristics including high light output, high gamma ray stopping efficiency, fast response, low cost, good proportionality, and minimal afterglow that the material is useful for gamma ray spectroscopy, medical imaging, nuclear and high energy physics research, diffraction, non-destructive testing, nuclear treaty verification and safeguards, and geological exploration. The timing resolution of the scintillators of the present invention provide compositions capable of resolving the position of an annihilation event within a portion of a human body cross-section.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional PatentApplication No. 60/505,636, filed Sep. 24, 2003, which is incorporatedherein by reference in its entirety.

STATEMENT AS TO RIGHTS TO INVENTIONS MADE UNDER FEDERALLY SPONSOREDRESEARCH OR DEVELOPMENT

Aspects of this research was conducted with funding provided by theDepartment of Energy and the National Institute of Health. Contract No.DE-AC03-76SF00098 and Grant Nos. R01-CA67911 and P01-HL25840,respectively. The United States Government may have certain rights inthe application.

BACKGROUND OF THE INVENTION

The present invention relates generally to scintillators. Morespecifically, the present invention provides lutetium iodide (LuI)scintillators for use with medical imaging scanner systems, such asgamma ray spectroscopy and time-of-flight positron emission tomography.

Scintillators are the most widely used detectors for spectroscopy ofenergetic photons (X-rays and gamma-rays). These detectors are commonlyused in nuclear and high energy physics research, medical imaging,diffraction, non-destructive testing, nuclear treaty verification andsafeguards, and geological exploration. Important requirements for thescintillation crystals used in these applications include high lightoutput, high gamma ray stopping efficiency (attenuation), fast response,low cost, good proportionality, and minimal afterglow. Theserequirements have not been met by any of the commercially availablescintillators, and there is continued interest in the search foradditional scintillators with enhanced performance.

One form of medical imaging is called positron emission tomography andis better known by its acronym PET. PET is a functional imagingtechnique with the potential to quantify the rates of biologicalprocesses in vivo. See J. T. Bushberg, J. A. Seibert, E. M. Leidholdt,and J. M. Boone, The Essential Physics of Medical Imaging, Williams andWilkins, (1994). The availability of short lived positron-emittingisotopes of carbon, nitrogen, oxygen and especially fluorine allowsvirtually any compound of biological interest to be labeled in traceamounts and introduced into the body for imaging with PET. Thedistribution of the tracer is imaged dynamically, allowing the rates ofbiological processes to be calculated using appropriate mathematicalmodels. PET imaging can provide diagnosis for symptoms of diseases suchas cancer, Alzheimer's disease, head trauma, and stroke. Phelps, M. E.Positron emission tomography provides molecular imaging of biologicalprocesses, Proc Natl Acad Sci, 97(16), 9226-9233, (2000).

In PET (or PET scan), the patient is injected with a molecule labeledwith a positron-emitting radioactive element. In some applications theradiotracer is distributed through the body, and concentrated in (orexcluded from) target tissues of interest. The radioactive materialdecays by emission of a positron, or antiparticle of thenegatively-charged electron. The positron is slowed down within a shortdistance from the emission point and forms a short-lived “atom”consisting of the positron and an electron from a nearby atom. The“atom,” referred to as positronium, decays by the annihilation of itsconstituents. This annihilation produces two essentially back to back511 keV gamma rays. When both of the gamma rays are detected bydetectors surrounding the body, it can be assumed with high probabilitythat the emission point was somewhere along a line joining the twodetectors. Without additional information, the probability that theradiotracer was located on any one point in the body that the detectionline intersects is equal for all points in the line, and hence, for allpoints in the body being scanned.

A variety of algorithms have been developed that make it possible toform an image from a collection of such lines. The quality of the imageimproves, in general, as the number of lines increases. Similarly, asthe signal-to-noise of the image depends on the square root of thenumber of lines, each line representing one annihilation event, morelines offer an improved signal to noise ratio. Nevertheless, a commonaspect of all image formation or reconstruction algorithms is that thenoise increases in the process of deciding where along the detected linethe annihilation event is likely to have occurred. In one aspect, thiseffect can be thought of in terms of energy and work: The detected linesrepresent the energy in the image and a large part of that energy isused up as work in localizing the annihilation along the particularline, instead of contributing to image quality. If the bodycross-section is 30 cm and the desired localization accuracy of anannihilation event is 5 mm, localization requires reducing theuncertainty of its location by a factor of 60.

The annihilation gamma-rays travel at a speed of about 30 cm/nanosecond(1 foot/ns). The timing accuracy of detectors currently usedcommercially in PET cameras is a few nanoseconds (ns). Timing resolutionis typically applied to two aspects of PET: One use is to reduceaccidentals (the overlapped detection of two unrelated gamma-rays). Theother use is in timing signals for localization purposes. While anyimprovement in time resolution aids in accidentals reduction, until thetime resolution drops substantially below about 1 ns, greater timeresolution will not help in localization of an annihilation event withina target of about 30 cm—the more common presentation in human-PETscanning. That is, if a body cross-section is about 30 cm, that veryfact localizes the event without any recourse to time resolution. Withthat limitation, an improvement in localization from 4 ns (4 feet) to 1ns (1 foot) offers no improvement to image quality. On the other hand, atimimg signal improvement from 1 ns to 500 picosecond (0.5 ns) reducesthe uncertainty of event location by a factor of 2. To appreciate thevalue of this particular improvement, it is to be noted that the factorof 2 increase in time resolution accuracy results in a correspondingincrease in signal-to-noise ratio. This results in the equivalent of afactor of 4 increase in detected annihilation events. Placed in adifferent context, under the same circumstances, an image (actually adata set) that may take 15 minutes to obtain with a 1 ns timeresolution, is obtained in under about 4 minutes when the timeresolution is 500 ps.

FIG. 1 illustrates the principle on which the location along thedetected line is used to improve image formation.

PET is playing a prominent and an increasingly visible role in modemresearch and clinical diagnosis. However, there is a need forimprovement in PET instrumentation in order to exploit the fullpotential of this promising technique. The performance of current PETsystems is limited by the available detector technology. Scintillationcrystals (herein referred to a “scintillators”) coupled tophotomultiplier tubes are commonly used as detectors in PET systems.Important requirements for the scintillators used in PET systems includefast response, high sensitivity, high light output, high energy andtiming resolution, and low cost. High energy resolution is importantbecause it allows rejection of scattered events. High timing resolutionis important because it allows rejection of random events. Furthermore,if sufficiently fast scintillators become available, time-of-flight(TOF) information could be utilized to obtain better event localizationcompared to conventional PET, which can lead to enhanced signal-to-noiseratio in the reconstructed image. Budinger TF. Time-of-flight positronemission tomography: status relative to conventional PET J. Nucl. Med.24: 73-78, (1983).

It is generally recognized that afast timing scintillator in PET cameraswill enable time-of-flight-PET when the timing accuracy and/or timingresolution is below 1 ns. Hitherto no true time-of-flight PET device hasbeen enabled. Barium fluoride (BaF₂), lutetium orthosilicate (LSO) andbismuth germnanate (BGO) have been suggested as potentially usefulscintillation materials, but none of these materials has the 500picosecond (ps) or less time resolution needed to achieve a successfuldevice. BGO, however, has a poor energy resolution and slow response,which limits its performance in 3D whole body imaging. The energyresolution of LSO is variable and is limited by its non-proportionality.Moses WW, Current Trends in Scintillator Detectors and Materials, Nucl.Inst. And Meth., A 487, p. 123-128, (2002). BaF₂ actually provides a˜250 ps (FWHM) timing resolution, but it has a low emission intensityfor the fast component and emits in blue region of the spectrum wherespecial photomultiplier tubes (PMTs) with quartz windows are requiredfor readout. It is noted that a number of plastic scintillators have atime resolution below 500 ps, but due to inadequate stopping power,(attenuation length at 500 keV is typically greater than 10 cm), thesescintillators are not suitable for medical uses.

The present invention provides a cerium doped rare-earth halidescintillator, lutetium iodide (Lu_(1-x)I₃:Ce_(x)). The crystals providea very fast scintillator materials capable of resolving the position ofan annihilation event within a portion of a human body cross-section(less than 400 ps). Specifically, the very fast scintillator materialcomprises LuI₃ doped with various concentrations of cerium. Crystals ofthis material have been grown and characterized and they providescintillators with properties suitable for many uses including use as agamma ray detector, in nuclear and particle physics, X-ray diffraction,non-destructive evaluation, treaty verification and non-proliferationmonitoring, environmental cleaning, geological exploration and medicalimaging. The timing resolution measured for the LuI₃:Ce crystals of thepresent invention demonstrate that the compositions provide ascintillator particularly useful in PET, including Time-of-Flight (TOF)PET devices and methods.

Attention is drawn to several references in the field, the teachings ofwhich are incorporated herein by reference (as are all references citedherein):

U.S. Pat. No. 6,362,479, “Scintillation detector array for encoding theenergy, position, and time coordinates of gamma ray interactions,”discloses a scintillator-encoding scheme that depends on thedifferential decay time of various scintillators. The use of lutetiumorthosilicate-lutetium orthosilicate (LSO-LSO) crystals with a timeresolution of 1.6 ns is also discussed. A time resolution of 1.6 ns isequivalent to an approximately 50 cm uncertainty, which is as large asthe cross-sectional dimension of the human body, and not useful inTOF-PET.

U.S. Pat. No. 5,453,623, “Positron emission tomography camera withquadrant-sharing photomultipliers and cross-coupled scintillatingcrystals.” Discloses arrangement of hardware elements in PET camera anduse of scintillators. Only specific scintillator disclosed is BGO.

Moses et al., “Prospects for Time-of-Flight PET using LSO Scintillator,”IEEE Trans. NucL. Sci. 46:474-478 (1999). Discloses measurements of thetiming properties of lutetium orthosilicate (LSO) scintillator crystalscoupled to a PMT and excited by 511 keV photons.

U.S. Pat. No. 5,319,203 and U.S. Pat. No. 5,134,293, both entitled“Scintillator material.” Discloses Cerium fluoride and thallium dopedCerium fluoride as “improved” scintillator material.

U.S. Pat. No. 5,039,858, “Divalent fluoride doped cerium fluoridescintillator.” Discloses additional doped cerium fluoride scintillators.

U.S. Pat. No. 4,510,394, “Material for scintillators.” Discloses bariumfluoride as scintillator material.

van Loef et al., “High energy resolution scintillator: Ce³⁺ activatedLaBr₃ ”, AppL. Phys. Lett. 79:1573-1575 (2001).

van Loef et al., “Scintillation properties of LaBr₃:Ce³⁺ crystals: fast,efficient and high-energy-resolution scintillators”, Nucl. Instr. Meth.Physics Res. A 486:254-258 (2002). Discloses certain characteristics ofcerium doped LaBr₃ compositions including, light yield, andscintillation decay curve. The rise time and time resolution of thecompositions are not disclosed or suggested.

WO 01/60945, “Scintillator crystals, method for making same, usethereof”, Discloses inorganic scintillator material of the generalcomposition M_(1-x)Ce_(x)Br₃, where M is selected from lanthanides orlanthanide mixtures of the group consisting of La, Gd, and Y. X is themolar rate of substitution of M with cerium, x being present in anamount of not less than 0.01 mol % and strictly less than 100 mol %. Therise time and time resolution of the various compositions are notdisclosed or suggested.

BRIEF SUMMARY OF THE INVENTION

The present invention provides a cerium doped lutetium iodide (LuI₃:Ce)scintillator. Crystals of LuI₃:Ce have been manufactured using meltbased Bridgman method. Scintillation properties of small LuI₃:Cecrystals (˜0.3 cm³, doped with Ce³⁺) include a peak emission wavelengthfor LuI₃:Ce at 474 nm which is well matched to PMTs as well as silicondiodes used in nuclear instrumentation.

The principal decay-time constant for LuI₃:Ce (with 5% Ce³⁺) is 25 ns,which is faster than the decay-time constant of commercial PETscintillators such as BGO, LSO, NaI:T1 and GSO. The light output ofLuI₃:Ce is ˜50,000 photons/MeV which is about 2 times higher than thatof LSO, about 6-7 times higher than that for BGO and GSO, and about 31%higher than that of NaI:T1.

The initial photon intensity—a figure of merit for timing applicationsis also higher for LuI₃:Ce compared to BGO, LSO, NaI:T1 and GSO. Thecombination of higher light output and faster response for LuI₃:Cecompared to existing PET scintillators promises high energy and timingresolution with LuI₃:Ce scintillators. These properties are veryattractive in whole body PET imaging where the ability to reject randomsand scatter needs to be improved. The measured timing resolution of aLuI₃:Ce crystal in coincidence with a BaF₂ crystal to be 210 ps (FWHM).Thus, Lul₃:Ce also provides the opportunity for time-of-flight (TOF) PETimaging which would provide additional gain in signal to noise ratio andimage quality. Due to its high atomic number constituents and highdensity (5.6 g/cm³), LuI₃:Ce provides high gamma ray sensitivity. Themean penetration depth of 511 keV photons in LuI₃:Ce is about 1.7 cm,which is comparable to that for GSO and slightly larger than that forLSO and BGO. The mean penetration depth of LuI₃:Ce is substantiallyshorter than that for NaI:T1.

Since LuI₃ melts congruently, it can be grown using crystal growthtechniques such as Bridgman and Czochralski which are generally easy toscale-up. Furthermore the melting point of LuI₃ is 1050° C., which issubstantially lower than the melting point of LSO and GSO (>2000° C.).As a result, the eventual cost of LuI₃:Ce can be expected to beconsiderably lower than that of LSO and GSO. This issue is particularlyrelevant in modem PET instrumentation where the high cost of thedetector components can be a major limitation. Thus, LuI₃:Ce appears tobe a very promising scintillator for PET imaging.

In one embodiment the present invention comprises a very fastscintillator comprising lutetium iodide and a trivalent cerium dopant.In one configuration, said dopant is present at about 0.1 % or more andless than or equal to about 100% by molar weight (e.g., a CeI₃scintillator), and particularly from about 0.5% to about 5% by molarweight, and more particularly about 5.0% by molar weight.

In certain embodiments the scintillator has a fast component with adecay constant of about 23 to about 31 nanoseconds. These embodimentsalso have a timing resolution of about 210 ps making the compositionsuseful for time-of-flight PET. Optionally, the scintillator may have aslow component with a decay constant of about 120 to about 230nanoseconds.

In another aspect this invention comprises a positron emission scannersystem comprising a patient area and an assembly of radiation detectorsdisposed adjacent the patient area. The radiation detectors comprise afast scintillator comprising lutetium iodide and a trivalent ceriumdopant. A scintillation light detector or photomultiplier tube areoptically coupled to the scintillator. A control system is coupled tothe light detectors or photomultiplier tube.

In one configuration, the dopant is present at about 0.1% or more andless than or equal to about 100% by molar weight, preferably betweenabout 0.5% or more and less than or equal to about 5.0% by molar weight,and most preferably about 5.0% by molar weight. The compositionscomprising lutetium iodide and a trivalent cerium dopant also have highlight output, sufficient stopping power and energy and/or timingresolution required for a positron emission scanner system useful fortime-of-flight measurements.

In some configurations, the scintillator is used in coincidencedetection positron emission tomography by recording the differentialarrival time of two photons so as to localize the annihilation event.Advantageously, the localization is carried out within a distance thatis less than about 30 cm.

The positron emission tomography scanner typically includes two or moreradiation detectors, in which each scintillation light detector of theradiation detector comprises a position sensitive detector or array. Thescanner typically includes means to correct for different timing offsetsof each of the individual radiation detectors. Such timing offsets o theindividual radiation detectors are stored in a memory in the controlsystem. For example, in one configuration, for each radiation detectorthe timing offsets are subtracted from each gamma-ray time arrival valueprior to computation of a localization. In another configuration, timingsignals of individual radiation detectors are equalized by anintroduction of individual hardwired delays in signal readoutelectronics in the control system.

Optionally, the scanner, comprising two or more scintillators, uses Cedoped LuI₃ in combination with other scintillators.

In a further embodiment the present invention comprises an X-raycomputed tomography (CT) scanner system comprising a patient area and apenetrating x-ray source. A detector assembly is positioned adjacent thepatient area on a substantially opposite side of the patient area. Thedetector assembly comprises a scintillator comprising lutetium iodideand a trivalent cerium dopant.

In one embodiment, the dopant is typically present at about 0.1% or moreand less than or equal to about 100% by molar weight, preferably betweenabout 0.5% or more and less than or equal to about 5.0% by molar weight,and most preferably about 5.0% by molar weight. The cerium dopedlutetium iodide fast scintillator also possess additionalcharacteristics necessary for an X-ray CT scanner system, such as forexample, high detection efficiency (high density and atomic number),high light output, linear light output with energy, fast decay time, lowcost and ease of crystal fabrication.

An additional embodiment of the present invention is a method ofperforming time-of-flight positron emission tomography. Such methods usea scintillator comprising lutetium iodide (LuI₃) and trivalent cerium asa dopant. The scintillator typically has a fast component with a decayconstant of about 23 to about 31 nanoseconds, and a time resolution ofless than 500 picoseconds (ps), and preferably below 0.4 nanoseconds(ns). The scintillator may comprise a slow component with a decayconstant of about 120 to about 230 nanoseconds.

The cerium dopant can be present at about 0.1% or more and less than orequal to about 100% by molar weight. More particularly, the ceriumdopant is typically present at about 0.5% or more, preferably betweenabout 0.5% and about 5.0% by molar weight, and most preferably at about5.0% by molar weight. The imaging method comprises injecting orotherwise administering a patient with a detectable label, and after asufficient period of time to allow localization or distribution of thelabel, placing the patient within the field of view of the device. Whena 511 keV gamma ray is detected by any one first detector, the deviceopens a time window (no less than up to about 1 ns long for the wholebody, but longer if the time resolution of the device is worse than 1ns, e.g., 10 ns for one of the scanners described above). If another 511keV event is detected within this time window at a second detector thatis across the body from the first detector (or, in some embodiments,where each detector comprises position sensing built within it), theposition within the detector and the detector's position are recorded,as well as the arrival times. Each positive pair defines a line. Fromthe known body size, the length of the line need not be the distancebetween detectors, it can be just the size of the cross the body fromthe first detector, this event is accepted as a coincidence. Theposition of the first and second detectors (or, in some configurationswhere each detector comprises position sensing built within it, theposition within the detector and the detector's position) are recorded,as well as the arrival times.

Each position pair defines a line. From the known body size, the lengthof the line needed not be the distance between detectors, it can be justthe size of the body cross-section. If there is no Time-of-Flight (TOF)information, equal probability is assigned to each point on the line.The reconstruction of the image then proceeds by one of the dozens ofalgorithms known in the art. If TOF information is available, then theprobability of origin of the event along the line can be represented asa Gaussian or similar distribution of width equal to the TOF FWHM,centered on the most probable point. Similar reconstruction algorithms,modified to take advantage of the TOF information can be used forreconstruction, and these modifications are also well known in the art.

In another embodiment, the present invention provides a method oflocalizing a positron annihilation event within a portion of a humanbody cross-section. In the method a positron emission tomography scanner(or camera) is used wherein the scanner comprises a scintillatorcomprising lutetium iodide (LuI₃) and trivalent cerium as a dopant.

The scintillator may have a fast component with a decay constant ofabout 23 to about 31 nanoseconds, a decay constant of about 120 to about230 nanoseconds, an attenuation length of about 1.7 cm, a light outputof about 47,000 to 50,000 photons/MeV, an initial photon intensity ofabout 1,200 photons/(ns×MeV), and a time resolution of about 0.210nanoseconds.

For a fuller understanding of the nature and advantages of the presentinvention, reference should be had to the ensuing detailed descriptiontaken in conjunction with the accompanying drawings. The drawingsrepresent embodiments of the present invention simply by way ofillustration. The invention is capable of modification in variousrespects without departing from the invention. Accordingly, the drawingsand description of these embodiments are illustrative in nature, and notrestrictive.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates the principle on which the location along thedetected line is used to improve image formation.

FIG. 2 represents output signals of two scintillators with the samelight output, but one with a rise time speed twice as fast as the other.

FIG. 3 shows the signal output from two scintillators of equal speed butdifferent light output.

FIG. 4 plots a radioluminescence spectrum of a LuI₃:Ce (0.5% Ceconcentration) scintillator. The main emissions at 474 and 515 nm aredue to Ce³⁺ luminescence.

FIG. 5 provides a time profile of LuI₃:Ce (5% Ce concentration)scintillation.

FIG. 6 presents the ¹³⁷Cs spectra recorded with LuI₃:Ce and BGOscintillators coupled to a PMT under substantially identical operatingconditions.

FIG. 7 plots the ¹³⁷Cs spectrum collected with a 0.3 cm³ LuI₃:Ce (5% Ceconcentration) crystal coupled to a PMT. The energy resolution of 662keV photopeak is approximately 10% (FWHM) at room temperature.

FIG. 8 plots the proportionality of response as a function of gamma rayenergy for a LuI₃:Ce (0.5% Ce concentration) scintillator.

FIG. 9 provides the coincidence timing resolution plot for BaF₂ andLuI₃:Ce (0.5% Ce concentration) scintillators upon irradiation with 511keV gamma ray pairs. The timing resolution is 210 ps (FWHM).

FIG. 10. shows a schematic of a positron emission scanner system.

FIG. 11 shows a schematic of the detector arrangement for a positronemission scanner system.

FIG. 12 shows a schematic of an x-ray computed tomography scanner systemencompassed by the present invention.

DETAILED DESCRIPTION OF THE INVENTION

This invention will be better understood with resort to the followingdefinitions:

A. Rise time, in reference to a scintillation crystal material, shallmean the speed with which its light output grows once a gamma ray hasbeen stopped in the crystal. The contribution of this characteristic ofa scintillator combined with the decay time contribute to a timingresolution. A timing resolution of less than 500 picosecond (ps) is ofparticular interest for use in methods comprising time-of-flightdetection of an annihilation event as originating within about a 30 cmdistance.

B. Fast timing scintillator should be capable of localizing anannihilation event as originating from within about a 30 cm distance,i.e., from within a human being scanned. This typically requires atiming resolution of about 500 ps or less.

C. Timing accuracy or resolution, usually defined by the full width halfmaximum (FWHM) of the time of arrival differences from a point source ofannihilation gamma-rays. Because of a number of factors, there is aspread of measured values of times of arrival, even when they are allequal. Usually they distribute along a bell-shaped or Gaussian curve.The FWHM is the width of the curve at a height that is half of the valueof the curve at its peak.

D. Light Output shall mean the number of light photons produced per unitenergy deposited by the detected gamma-ray, typically the number oflight photons/MeV.

E. Stopping power or attenuation shall mean the range of the incomingX-ray or gamma-ray in the scintillation crystal material. Theattenuation length, in this case, is the length of crystal materialneeded to reduce the incoming beam flux to 1/e.

F. Proportionality of response (or linearity). For some applications(such as CT scanning) it is desirable that the light output besubstantially proportional to the deposited energy.

G. Coincidence timing window/coincidence detection shall mean the lengthof time allowed for deciding whether two detected 511 keV gamma-raysbelong to the same positron annihilation event. This window is desiredto be as short as possible, but no shorter than the time it takes thegamma-rays to travel through the body (>1 nsec).

H. Single line time-of-flight (TOF) localization shall mean the processby which, through timing of the signals, the position of theannihilation event is localized to a portion of the line joining thedetectors, this portion being smaller than the length of the line.

I. Position sensitive detector or array shall mean a detector where theposition of the gamma-ray interaction within the detector is determined.In some embodiments this is done through the Anger principle of lightdivision (well known in the state of the art). For instance, there canbe a photodetector at each end of the crystal and the proportion oflight reaching each detector determines position, or an array ofphotodetectors where the center of mass of the light distributiondetermines position (i.e., the closest detectors get more light).

J. Method to correct for different timing offsets of an individualdetector shall be understood to include, among others, software codethat stores each detector's individual timing delay and code to subtractfrom each timing signal this pre-stored value. Method to introducethrough delay lines (cables through which the signal travels) a fixeddelay for each detector, so that their signals all have the same arrivaldelay at the timing electronics.

A property of a scintillator crystal material is the speed with whichits light output grows once a gamma ray has been stopped in the crystal.This property is characterized by the rise time of the scintillatorcrystal. An example is shown in FIG. 2 for two scintillators with thesame light output, but one with a rise time speed twice as fast as theother. There is a noise level (due to readout electronics) that does notallow the signal to be reliably detected until it exceeds a certainthreshold (70 in this example). Both signals start at point 10 on thehorizontal axis of the graph, and the faster scintillator crosses thethreshold above noise faster. Consequently, variations in timing fromdifferent pulse strengths will be smaller for the faster rise time speedscintillator. The faster rise time scintillator permits a higher timeresolution.

Increased or high light output impacts the signal-to-noise ratio inscintillation detection. Given the noise generally inherent in thereadout electronics, higher light output leads to better energyresolution. Better energy resolution is useful in identifying andexcluding gamma-ray detections of gamma-rays that have scattered in thebody yielding a “false” line as compared with those that have notscattered. Higher light output also enables improved accuracy in timing.As the signal rises towards a higher peak value, it crosses anoise-dictated threshold of detectability sooner. As a result variationsin signal output (due to finite energy resolution) lead to a smallerrange of time differences in crossing the threshold. FIG. 3 shows thesignal output from two scintillators of equal speed (rise time) butdifferent light output.

Stopping power is an aspect of detection efficiency. Stopping power,particularly at 511 keV, is an important parameter for a scintillatormaterial for use in a PET scanner or camera. This efficiency isdependent, in part, on the density and average atomic number of thescintillator material. High values of both density and average atomicnumber tend to increase detection or stopping power of the scintillator.A high stopping power is advantageous, and the higher, the better. Thehigh attenuation power of LuI₃ (short attenuation length), means thatphysically smaller detectors can be built while maintaining gooddetection efficiency. Smaller detectors are understood by those familiarwith the art as providing better time and spatial resolution.

In the practice of the present invention, attention is paid to thephysical properties of the scintillator material. In most embodiments arobust scintillator crystal or ceramic is preferred. Similarly, inparticular applications, properties such as hygroscopy (tendency toabsorb water), brittleness (tendency to crack), and crumbliness shouldbe minimal.

Table I below presents properties of two conventional positron scannersor cameras currently in the market. The time resolution of one of them,the TOFPET TTV 03, at 650 ps does not significantly localize anannihilation event within the typical 30 cm cross-section of a human.For such a time resolution, up to 40% of detected events can belocalized to within 10 cm, and approximately 15% will appear to arisefrom outside a 30 cm body cross-section. A time resolution of 650 ps isnot acceptable for use in PET TOF localization. A time resolution ofless than 500 ps is required. TABLE I TOFPET PET TTV 03 Siemens/CTI Ringdiameter [mm] 890 820 Number of rings 4-6  24 Number of detectors per324 784 ring Crystal dimension (mm) 7 × 18 × 45 2.9 × 5.9 × 30 Type ofcrystal BaF₂ BGO Spatial resolution [mm]  5  4 Time resolution (ps) 650750

The present invention includes a method of appropriately doping LuI₃with trivalent Ce, to obtain a material capable of high light output(greater than about 50,000 photons/MeV at room temperature) well matchedto photo-detection (FIG. 4), fast response (FIGS. 5 and 9) and totime-of-flight positron detection localization capabilities (timeresolution of less than 500 ps; FIG. 9). LuI₃ doped with Ce³⁺ atconcentration of about >0. 1% Ce³⁺ molar weight, and particularlyconcentrations between about 0.5% and about 5.0%, and up to 100%, havebeen found to be useful in medical imaging including PET andtime-of-flight positron detection localization, and the like. TABLE IIProperties of Scintillators Light Attenuation Initial Output WavelengthLength Photon Principal (Photons/ of Emission (511 keV) Intensity DecayMaterial MeV) (nm) (cm) (Photons/(ns × MeV)) Time (ns) NaI(Tl) 38,000415 3.3   165  230 CsI(Tl) 52,000 540 1.9   50 1000 LSO  24000 420 1.2  600  40 BGO  8,200 505 1.1   30  300 BaF₂ 10,000 310 slow 2.3 3,400(total)  620 slow ˜2,000 fast 220 fast   0.6 fast GSO  7,600 430 1.5  125  60 CdWO₄ 15,000 480 1.1    3 5000 YAP 20,000 370 2.1   570  26LaBr₃ (0.5% Ce)¹ 61,000 360 2.1 1,850  31 LaBr₃ (0.5% Ce)² 68,000 3702.1 2,600  26 LaBr₃ (5% Ce)² 62,500 370 2.1 4,300  16 LaBr₃ (10% Ce)²64,500 370 2.1 3,900  16 LaBr₃ (20% Ce)² 64,000 375 2.1 3,600  17 LaBr₃(30% Ce)² 69,500 375 2.1 3,650  18.6 LuI₃ (0.5% Ce) 47,000 470 1.7 1,300 31 LuI₃ (5% Ce) 50,000 474 1.7 1,800  25¹Data based on the Delft University of Technology results (See, van Loefet al., Nucl. Inst. Meth. Phys. Res. A 486: 254-258 (2002)).²Data and results provided in US Patent Application No. unassigned,filed; Attorney Docket No. 22071-000110US.

Compared to CsI, which is among the scintillation materials with thehighest known light output, LuI₃ produces about the same amount of light(approximately 47,000 to about 50,000 photon/MeV), a fast principaldecay constant (about 23 to about 31 ns), has a slightly shorterattenuation length, a very fast light output (initial photons), and theenergy resolution of LuI₃:Ce coupled to a PMT and measured at 662 keVwas about 10% (FWHM). Timing resolution of a LuI₃-PMT operating incoincidence mode was measured to be about 210 ps (FWHM).

LuI₃ has hexagonal crystal structure, a density of 5.6 g/cm³, and can begrown directly from the melt by techniques such as Bridgman andCzochralski. This is a useful property because these melt-basedtechniques are generally easier to scale-up and are used in commercialproduction of scintillators. Crystals have been usefully grown usingthese methods although other methods for their growth are well known tothe skilled artisan. LuI₃ is moisture-sensitive and therefore should behermetically packaged to prevent exposure to moisture.

As will be understood by one of skill in the art, fast scintillators areused in conjunction with methods to calibrate each detector to correctfor differential time lags that confuse relative timing measurements. Inparticular embodiments, such corrections are performed by introducinghardwired delays of appropriate lengths or by software processing basedon pre-stored delay times for each detector. Within the practice of thepresent invention scintillators are used in individual detectors(detector modules) or read by position-sensitive photodetectors orarrays of photodetectors that detect the light from the scintillation ofthe crystal or ceramic.

The applications of these fast detectors are not limited to PET cameras.They are also useful in applications where fast decay of the lightsignal is desirable. One such application is X-ray computed tomography(CT), where, as rotation times and individual detector size decreases,detector response time become more important. The high linearity ofoutput of the scintillators of the present invention is of particularuse in this application.

Notable parameters for the scintillation crystals used in spectroscopyof energetic photons (gamma rays) as well as neutrons at roomtemperature applications include high light output, high stoppingefficiency, fast response, low cost, good linearity, and minimalafterglow.

EXAMPLE 1

LuI₃ Crystals, Bridgman Method

In making crystals, ultra-dry forms of LuI₃ and CeI₃ were used with99.99% purity. A two zone Bridgman furnace was used with temperature inthe hotter zone above the melting point LuI₃ (1050° C.) and that of thecooler zone below the melting point. The amount of CeI₃ in the feedmaterial was adjusted to produce LuI₃ samples with varying Ce³⁺concentration. Most growth runs were performed with 0.5% and 5.0% (on amolar basis) cerium concentration, although runs can also be performedwith other Ce³⁺ concentrations (e.g., 0.1%, 10%, 20%, 30%, 40%, 50%,60%, and up to or less than 100%). LuI₃ crystals as large as ˜1 cm³ weregrown using this process. These crystals were cut from the solid ingotsto produce small samples (≦0.3 cm³ size) for measurements.

Scintillation properties of the LuI₃ crystals were then measured. Forsome measurements, packaged samples were used because LuI₃:Ce crystalswere sensitive to moisture. This involved placing a LuI₃:Ce crystalinside a metal can on a quartz window. The crystal was attached to thequartz with a clear optical epoxy (e.g., EPO-TEK #301-2) The spacearound the crystal in the can was filled with SiO₂ powder. The top ofthe can was finally sealed by attaching a metal disc using epoxy.

EXAMPLE 2

LuI:Ce, Scintillation Properties

Characterization of the scintillation properties of LuI₃ crystals grownby the Bridgman process involved measurement of the light output, theemission spectrum, and the fluorescent decay time of the crystals.Energy and timing resolution of the LuI₃:Ce crystals and theirproportionality were also measured. Based on its high atomic numberconstituents and high density, Lu_(1-x)I₃Ce_(x) show high gamma raystopping efficiency. The attenuation length of 511 keV photons inLuI₃:Ce was 1.7 cm.

Emission Spectrum

The emission spectrum of the LuI₃:Ce scintillators was measured. TheLuI₃:Ce samples were excited with radiation from a Philips X-ray tubehaving a copper target, with power settings of 35 kVp and 15 mA. Thescintillation light was passed through a McPherson monochromator anddetected by a Hamamatsu R2059 photomultiplier tube with a quartz window.

FIG. 4 shows the normalized emission spectra for LuI₃:Ce samples with a0.5% Ce concentration. As seen in FIG. 4, the peak emission wavelengthfor the LuI₃:Ce sample is at about 474 nm. The smaller peaks observed inthe 550-620 nm region may be due to impurities that may be present inthe Sample. Emission spectrum measured for LuI₃:Ce with 5.0% Ce was verysimilar to that displayed in FIG. 4. The peak emission wavelength of 474nm for LuI₃:Ce matches well with the spectral response of thephotomultiplier tubes as well as silicon photodiodes that are used inscintillation detection. The emission spectrum for LuI₃:Ce with 5.0%Ce³⁺ is similar to that shown in FIG. 4.

Decay Time

The fluorescent decay time profile of a LuI₃:Ce scintillation has beenmeasured using the delayed coincidence method as described in Bollingerand Thomas (Rev. Sci. Instr., 32:1044 (1961), the teachings of which areincorporated herein by reference) by exposing the crystal to 662 keVgamma rays (¹³⁷Cs source). FIG. 5 shows the time profile recorded for aLuI₃:Ce (5.0% Ce concentration) sample along with a multi-exponentialfit to the data. The principal decay constant for the sample is about 25ns (most likely arising from direct electron-hole capture on Ce³⁺ site)and this component covers about 88% of the integrated light output. Alonger decay component with 150 ns time constant is also present in theLuI₃:Ce sample with 5.0% Ce doping and covers about 12% of theintegrated light emission. For LuI₃ with 0.5% Ce doping, the principaldecay constant is about 31 ns, which covers >80% of the integrated lightoutput, with the remaining light emitted via a 230 ns decay component.Virtually no rise time is observed for the LuI₃ sample with 5.0% Ce³⁺. Arise time of ˜4nss was observed for the LuI₃ sample with 0.5% Ce³⁺. Fromthis, Applicants believe that a higher concentration of Ce will improvethe timing properties of the LuI₃:Ce crystals.

Light Output and Energy Resolution

The light output (or luminosity) of LuI₃:Ce crystals was measured withsamples (doped with 0.5% and 5.0% Ce³⁺ concentration) by comparing theirresponse and the response of a calibrated BGO scintillator (with 7500photons/MeV light output) to the same isotope (662 keV γ-rays, ¹³⁷Cssource, see FIG. 6). These measurements involved optical coupling of aLuI₃:Ce sample to a photomultiplier tube, irradiating the scintillatorwith 662 keV photons and recording the resulting pulse height spectrumusing standard NIM electronics. In order to maximize light collection,LuI₃:Ce crystals were wrapped in reflective, white Teflon tape on allfaces (except the one coupled to a photomultiplier (PMT)). FIG. 6 showsthe pulse height spectra for both LaBr₃:Ce and BGO under ¹³⁷Csirradiation and amplifier shaping time of 4.0 μsec. This shaping timewas long enough to allow full light collection from both thescintillators. The photomultiplier (PMT) bias and amplifier gain werethe same for both spectra. Based on the recorded photopeak positions forLuI₃:Ce and BGO, and by taking into account the photocathode quantumefficiency for BGO and LuI₃:Ce, the light output of LaBr₃:Ce crystalwith 5.0% Ce concentration was found to be about 50,000 photons/MeV atroom temperature, which is about 7 times higher than that of BGO (andabout 2 times higher than that of LSO). The light output measured for a0.5% Ce sample was similar to that measured for the 5.0% Ce sample. Themeasure light output of LuI₃:Ce is about 30% higher compared to that ofNaI(T1). Based on the light output and the decay time measurements theinitial photon intensity, a figure of merit in timing applications, wasestimated to be 1800 photon/(ns×MeV) for LuI₃:Ce (5.0% Ceconcentration), which is almost 10 times higher than that for NaI(T1).

Gamma Ray Spectroscopy

Gamma ray spectroscopy was conducted using a LuI₃:Ce scintillatorcoupled to a PMT. The scintillator was covered with Teflon tape andirradiated with 662 keV gamma rays (¹³⁷Cs source). The PMT signal wasprocessed with a spectroscopy amplifier (Canberra Model 2022) and agamma ray spectrum was collected using a multi-channel analyzer residingin a personal computer. FIG. 7 shows a ¹³⁷Cs spectrum collected in thismanner and the energy resolution of the 662 keV peak was measured to beabout 10% (FWHM) at room temperature. The energy resolution was mostlylimited by the optical quality of the LuI₃:Ce crystals available atpresent and significant improvement is expected as the crystal growth ofLuI₃:Ce is optimized and larger, higher quality crystals are produced.The low energy shoulder on the 662 keV photopeak has been attributed toescape of Lu K-edge X-rays.

Proportionality of Response

Proportionality of response (or linearity) of LuI₃:Ce scintillators wasalso evaluated. Non-proportionality (as a function of energy) in lightyield can be one of the important reasons for degradation in energyresolution of established scintillators such as NaI(T1) and CsI(T1)(Dorenbos et al., IEEE Trans. Nucl. Sci. 42:2190-2202 (1995); Moses,Nucl. Inst. Meth. A 487:123-128 (2002)). Light output of LuI₃:Ce wasmeasured under excitation from isotopes such as ²⁴¹Am (60 keV γ-rays),⁵⁷Co (14 keV X-rays and 122 keV γ-rays), and ¹³⁷Cs (662 keV γ-rays). ALuI₃:Ce crystal (0.5% Ce concentration) was wrapped in Teflon tape andcoupled to a PMT. Pulse height measurements were performed usingstandard NIM equipment with the scintillator exposed to differentisotopes. The same settings were used for PMT and pulse processingelectronics for each isotope. From the measured peak position and theknown γ-ray energy for each isotope, the light output (in photons/MeV)at each γ-ray energy was estimated. The data points were then normalizedwith respect to the light output value at 662 keV energy and the results(shown in FIG. 8) indicate that LuI₃:Ce is a fairly proportionalscintillator. Over the measured energy range, the non-proportionality inlight yield was about 5% for LuI₃:Ce, which is better than that for manyestablished scintillators. For example, over the same energy range, thenon-proportionality is about 35% for LSO:Ce and about 20% for NaI(T1)and CsI(T1) has been reported (Gillot-Noel et al., IEEE Trans. Nucl.Sci. 46:1274-1284 (1999)).

Coincidence Timing Resolution

Coincidence timing resolution of LuI₃:Ce crystals was measured. Thisinvolved irradiating a BaF₂ and LuI₃:Ce (0.5% Ce concentration)scintillators, each coupled to a fast PMT (Hamamatsu R-5320, operated at-2000V) with 511 keV positron annihilation γ-ray pairs (emitted by a⁶⁸Ga source). The BaF₂-PMT detector formed a “start” channel in thetiming circuit, while the LuI₃-PMT detector formed the “stop” channel.The signal from each detector was processed using two channels of aTennelec TC-454 CFD that had been modified for use with fast (sub-ns)rise-time PMTs. The time difference between the start and stop signalswas digitized with a Tennelec TC-862 TAC and a 16-bit ADC, resulting ina TDC with 7.5 ps per bin resolution. Data were accumulated until thecoincidence timing distribution had more than 10,000 counts in themaximum bin. FIG. 9 shows the coincidence timing resolution plotacquired in this manner with a LuI₃:Ce crystal having 0.5% Ce³⁺concentration. The timing resolution for this crystal was measured to beabout 210 ps (FWHM). The timing resolution for two BaF₂ detectors incoincidence with each other was measured to be 210 ps (FWHM) in thisstudy.

Basic PET Configuration

A PET camera system typically comprises of a polygonal or circular ringof radiation detectors (10) placed around a patient area (11), as shownin FIG. 10. In some embodiments radiation detection begins by injectingor otherwise administering isotopes with short half-lives into apatient's body placeable within the patient area (11). As noted above,the isotopes are taken up by target areas within the body, the isotopeemitting positrons that are detected when they generate pairedcoincident gamma-rays. The annihilation gamma-rays move in oppositedirections, leave the body and strike the ring of radiation detectors(10).

As shown in FIG. 11, the ring of detectors (10) includes an inner ringof scintillators (12) and an attached ring of light detectors orphotomultiplier tubes (14). The scintillators (12 ) respond to theincidence of gamma rays by emitting a flash of light (scintillation)that is then converted into electronic signals by a correspondingadjacent photomultiplier tube or light detectors (14). A computercontrol system (not shown) records the location of each flash and thenplots the source of radiation within the patient's body. The dataarising from this process is usefully translated into a PET scan imagesuch as on a PET camera monitor by means known to those in the art.

This invention has been discussed in terms of LuI₃:Ce crystalscintillators for use in PET, and particularly useful in time-of-flightPET. Such application of the technology is illustrative only. Many,indeed most, ionizing radiation applications will benefit from theinventions disclosed herein. Specific mention is made to X-ray CT, X-rayfluoroscopy, X-ray cameras (such as for security uses), and the like. ACT scanner as shown in FIG. 12, as well known to the skilled artisan,typically comprises a patient bed 22, a penetrating X-ray source 26(i.e., an X-ray tube), a detector assembly 29 and associated processingelectronics 29, and a computer and software for image reconstruction,display, manipulation, post-acquisition calculations, storage andretrieval 28.

1. A scintillator comprising lutetium iodide and a trivalent ceriumdopant
 2. The scintillator of claim 1, wherein said dopant is present atabout 0.1% or more and less than or equal to about 100% by molar weight.3. The scintillator of claim 1, wherein said dopant is present in anamount of from about 0.5% to about 5.0% by molar weight.
 4. Thescintillator of claim 3, wherein said dopant is present in an amount ofabout 5.0% by molar weight.
 5. The scintillator of claim 1 having a fastcomponent with a decay constant of about 23 to about 31 nanoseconds anda slow component, if present, with a decay constant of about 120 toabout 230 nanoseconds.
 6. A positron emission tomography scanner systemcomprising: a patient area; an assembly of radiation detectors disposedadjacent the patient area, wherein the radiation detectors comprise: ascintillator comprising lutetium iodide and a trivalent cerium dopant; ascintillation light detector or photomultiplier tube optically coupledto the scintillator; and a control system coupled to the light detectorsor photomultiplier tubes.
 7. The positron emission tomography scannersystem of claim 6, wherein said dopant is present at about 0.1% or moreand less than or equal to about 100% by molar weight.
 8. The positronemission tomography scanner system of claim 6, wherein said dopant ispresent at about 0.5% or more and less than about 5.0% by molar weight.9. The positron emission tomography scanner system of claim 6, whereinsaid dopant is present at about 5.0% by molar weight.
 10. The positronemission tomography scanner of claim 6, wherein said scintillator isused in coincidence detection positron emission tomography by recordingthe differential arrival time of two photons so as to localize theannihilation event.
 11. The positron emission tomography scanner ofclaim 10, wherein the localization is to within a distance of less thanabout 30 cm.
 12. The positron emission tomography scanner of claim 6,wherein the radiation detector comprises two or more radiationdetectors, wherein each scintillation light detector comprises aposition sensitive detector or array.
 13. The positron emissiontomography scanner of claim 12, further comprising a means to correctfor different timing offsets of each individual radiation detector. 14.The positron emission tomography scanner of claim 13, wherein timingoffsets of individual radiation detectors are stored in a memory in thecontrol system.
 15. The positron emission tomography camera of claim 14,wherein for each radiation detector the timing offsets are subtractedfrom each gamma-ray time arrival value prior to computation of alocalization.
 16. The positron emission tomography camera of claim 12,wherein timing signals of individual radiation detectors are equalizedby an introduction of individual hardwired delays in signal readoutelectronics in the control system.
 17. An X-ray computed tomographyscanner system comprising: a patient area; a penetrating X-ray source;and a detector assembly comprising a scintillator comprising lutetiumiodide and a trivalent cerium dopant wherein said dopant is present atabout
 0. 1% or more and less than or equal to about 100% by molarweight.
 18. The X-ray computed tomography scanner system of claim 17wherein said dopant is present at about 0.5% or more and less than about5.0% by molar weight.
 19. The positron emission tomography scannersystem of claim 17 wherein said dopant is present at about 5.0% by molarweight.
 20. The positron emission tomography scanner system of claim 17wherein said dopant is present at about 0.5% by molar weight.
 21. Amethod of performing time-of-flight-positron emission tomographycomprising: administering a patient with a detectable label; positioningthe patient within a field of view of a scintillator to detect emissionsfrom the patient, wherein the scintillator comprises lutetium iodide(LuI₃) and trivalent cerium as a dopant. 22 The method of claim 21wherein said scintillator has a timing resolution of less than 500 ps.23. The method of claim 21, wherein said Cerium dopant is present atabout 0.1% or more and less than or equal to about 100% by molar weight.24. The method of claim 21, wherein said Cerium dopant is present atabout 0.5% or more.
 25. The method of claim 21, wherein said Ceriumdopant is present at about 5.0% by molar weight.
 26. A method oflocalizing a positron annihilation event within a portion of a humanbody cross-section which comprises using a positron emission tomographyscanner system comprising a scintillator comprising lutetium iodide(LuI₃) and trivalent cerium as a dopant.